The present invention relates to the magnetic resonance arts. It finds particular application in conjunction with medical magnetic resonance imaging systems and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in conjunction with other types of magnetic resonance imaging systems, magnetic resonance spectroscopy systems, and the like.
In magnetic resonance imaging, a strong uniform static magnetic field B.sub.0 is generated, often by a superconducting magnet. This static magnetic field B.sub.0 polarizes the nuclear spin system of an object to be imaged. Superconducting magnets are commonly wound on a cylindrical body former mounted in an annular helium vessel surrounded by an annular vacuum vessel for thermal isolation. The superconducting magnet generates the static magnetic field, B.sub.0 along its own longitudinal axis and the common longitudinal axis of the cylindrical bore of the vacuum vessel, commonly denoted as the "z-axis".
To generate a magnetic resonance signal, the polarized spin system is first excited by a radio frequency magnetic field perpendicular to the z-axis. This RF field, denoted B.sub.1, is produced by an RF coil located inside the bore of the magnet and closely conforming thereto to maximize the space available to receive a patient. The RF magnetic field, which is changing in time in a sinusoidal waveform, is turned on and off to create short RF pulses to excite magnetization in the polarized object in the bore. More specifically, the RF pulses tip the magnetization out of alignment with the z-axis and cause its macroscopic magnetic moment vector to precess around the z-axis. The precessing magnetic moment, in turn, generates a radio frequency magnetic resonance signal that is received by the RF coil in a reception mode.
In magnetic resonance imaging, it is advantageous for the RF coil to have high sensitivity, high RF power efficiency, and a high signal-to-noise ratio. Also, the B.sub.1 magnetic field which it generates should be uniform. The sensitivity of the RF coil is defined as the magnetic field B.sub.1 created by a unit current. The signal-to-noise ratio is proportional to the sensitivity and to the square root of the coil quality factor Q.
To encode a sample spatially, magnetic field gradient pulses are applied after the RF excitation. The gradient magnetic fields are also applied in pulses to generate magnetic fields pointing in the z-axis, but changing in magnitude linearly in x, y, or z-directions. These gradient pulses are commonly denoted as G.sub.x, G.sub.y, and G.sub.z pulses, respectively. The gradient magnetic fields are generated by gradient magnetic field coils which are also located inside the magnet bore. Commonly, the gradient field coils are mounted on a cylindrical former around an outer periphery of the RF coil.
The whole body RF and gradient field coils have a sufficient inner diameter to receive the entire body of a patient within their circular bore. In order to receive a patient's body axially therethrough, the cylindrical whole body coils are relatively large in diameter, e.g. 60 cm. This large diameter tends to place the RF coils a significant distance from individual organs or small regions of the chest cavity to be examined. To overcome the signal-to-noise ratio and sensitivity problems inherent in this large spatial separation, surface coils are often used to receive magnetic resonance signals for individual organ examinations.
In order to optimize gradient performance, it is advantageous to place the gradient coils as close to the patient, hence as close to the RF coil, as possible without limiting patient accessible volume.
One of the problems in conventional magnetic resonance systems is that the radio frequency coil tends to couple with the gradient field coils. That is, the signals from the radio frequency coils induce analogous currents in the gradient coils. To eliminate this coupling, a shield coil is normally inserted between the gradient coil and the RF coil.
However, shield coils reduce the sensitivity of the RF coil. More specifically, secondary RF currents are induced in the shield during RF transmission. The induced secondary currents not only generate RF signals that tend to cancel the B.sub.1 field inside the coil, but also consume RF energy due to shield resistance. This lowers the unloaded coil Q factor. The secondary currents in the RF shield degrade the homogeneity of the primary B.sub.1 field generated by the RF coil. The closer the RF coil is to its shield, the more the coil sensitivity is reduced.
In order to counteract the negative effects of the shield coil, more RF power is supplied to the RF coil. Supplying more RF power requires larger, more costly power amplifiers and power transmission subsystems.
In order to achieve faster imaging, such as echo planar imaging, larger magnetic field gradients, which are applied for shorter periods of time, are needed. For example, echo planar imaging typically requires gradient magnetic fields on the order of 40 milliTesla per meter in pulses of 80 microsecond duration. This much higher spatial magnetic energy density, as compared to standard magnetic resonance imaging techniques, requires either more power to the gradient coil, or smaller gradient coils. Reducing the diameter of the gradient coil, of course, moves it closer to the RF coil and moves the RF shield therebetween closer to the RF coil. Again, placing the RF shield closer to the RF coil reduces the RF coil sensitivity, its power efficiency, and its signal-to-noise ratio. The power requirements for performing echo planar and other fast scanning techniques is a major impediment in adapting existing magnetic resonance scanner apparatus to perform fast scanning.
Another disadvantage of magnetic resonance imaging apparatus which incorporates RF shields resides in the high cost of such shields. The RF shields are typically etched from double-sided copper laminates which have low loss, high dielectric constant substrates. These high dielectric constant substrate laminates cost a few thousand dollars per MRI system.
The present invention contemplates a new and improved magnetic resonance apparatus which overcomes the above-referenced disadvantages by eliminating RF shielding coils between the RF and gradient coils.